The present invention relates to a hearing aid with an adaptive filter for suppression of acoustic feedback in the hearing aid.
It is well known in the art of hearing aids that acoustic feedback may lead to generation of undesired acoustic signals which can be heard by the user of a hearing aid.
Acoustic feedback occurs when the input transducer of a hearing aid receives and detects the acoustic output signal generated by the output transducer. Amplification of the detected signal may lead to generation of a stronger acoustic output signal and eventually the hearing aid may oscillate.
It is well known to include an adaptive filter in the hearing aid to compensate for acoustic feedback. The adaptive filter estimates the transfer function from output to input of the hearing aid including the acoustic propagation path from the output transducer to the input transducer. The input of the adaptive filter is connected to the output of the hearing aid and the output signal of the adaptive filter is subtracted from the input transducer signal to compensate for the acoustic feedback. A hearing aid of this type is disclosed in U.S. Pat. No. 5,402,496.
In such a system, the adaptive filter operates to remove correlation from the input signal, however, signals representing speech and music are signals with significant auto-correlation. Thus, the adaptive filter cannot be allowed to adapt too quickly since removal of correlation from signals representing speech and music will distort the signals, and such distortion is of course undesired. Therefore, the convergence rate of adaptive filters in known hearing aids is a compromise between a desired high convergence rate that is able to cope with sudden changes in the acoustic environment and a desired low convergence rate that ensures that signals representing speech and music remain undistorted.
The lack of speed of adaptation may still lead to generation of undesired acoustic signals due to acoustic feedback. Generation of undesired acoustic signals is most likely to occur at frequencies with a high feedback loop gain. The loop gain is the attenuation in the acoustic feedback path multiplied by the gain of the hearing aid from input to output.
Acoustic feedback is an important problem in known CIC hearing aids (CIC=complete in the canal) with a vent opening since the vent opening and the short distance between the output and the input transducers of the hearing aid lead to a low attenuation in the acoustic feedback path from the output transducer to the input transducer, and the short delay time maintains correlation in the signal.
Various measures are well known in the art to cope with acoustic feedback. For example, it is well known to keep the loop gain below a certain limit in order to prevent generation of feedback resonance. It is also known to adjust the phase of the feedback signal, to perform a frequency transpose, and to compensate for the feedback signal.
Typically, the acoustic environment of the hearing aid changes over time, and often changes rapidly over time, in such a way that propagation of sound from the output transducer of the hearing aid to its input transducer changes drastically. For example, such changes may be caused by changes in position of the user in a room, e.g. from a free field position in the middle of the room to a position close to a wall that reflects sound. Changes may also be generated if the user yawns or if the user puts the receiver of a telephone to the ear. Such changes, some of which may be almost instantaneous, are known to involve changes in attenuation of the feedback path of more than 20 dB.
It is known to keep the loop gain below a safe limit by limiting the gain adjustment in the hearing aid to a maximum allowable gain based on experience. However, a large safety margin is needed to cope with the above-mentioned variations in the acoustic environment and with variations in physical fitting of the hearing aid to the wearer. It is also known to determine the maximum allowable gain during fitting of the hearing aid to a specific user. However, a large safety margin is still needed. The safety margin prevents the capabilities of the hearing aid to be fully exploited, such as in situations where the gain could be adjusted to a value that is higher than the maximum allowable gain without generation of undesired sounds.
In order to be able to compensate for a severe hearing deficiency, it is desirable to be able to set a high gain in the hearing aid. However, the risk of generating oscillation, also denoted feedback resonance, restricts the maximum gain that may be employed, even in situations with a high attenuation in the acoustic feedback path.
In DE-A-19802568 and U.S. Pat. No. 5,016,280, a hearing aid is disclosed including a measuring system for determining the characteristics of the acoustic feedback path. A test signal is transmitted through the system in order to determine the characteristics of the feedback path.
In DE-A-19802568 the coefficients in a digital filter are determined based on the impulse response of the feedback path, and in U.S. Pat. No. 5,016,280 the filter coefficients of an adaptive compensation filter are calculated using a leaky LMS algorithm operating on white-noise signals transmitted through the feedback path.
The respective measuring systems are rather complicated and the duration of the determination is relatively long, and the normal function of the hearing aid is interrupted during the determination. Thus, the determination is performed at certain occasions only, e.g. when the user switches the hearing aid on. Thus, still, a relatively high safety margin for the gain is needed to cope with changes in the acoustic environment between determinations.
In U.S. Pat. No. 5,619,580 a hearing aid with an adaptive filter and a continuously operating measuring system is disclosed. A pseudo random noise signal is injected into the output signal. A monitoring system controls the gain of the hearing aid so that the loop-gain is kept below a constant value which may be frequency dependent. The filter coefficients of the adaptive filter are monitored and their update rate is adjusted according to a statistical analysis which complicates the system. It is another disadvantage of the system that a noise generator is needed and that the generated noise signal is always present. Moreover, the system increases the adaptation rate and thus deteriorates the signal quality when a change in acoustic environment is detected also in situations where the hearing aid is not operating close to resonance.
Thus, there is a need for an improved hearing aid that overcomes the above-mentioned disadvantages and substantially eliminates the requirement of a gain safety margin so that the operating gain in certain acoustic environments can be higher than for known hearing aids.
According to a first aspect of the invention, these and other objects are fulfilled by a hearing aid with an adaptive filter for compensation of acoustic feedback. The adaptive filter operates to estimate the transfer function from output to input of the hearing aid including the acoustic propagation path from the output transducer to the input transducer. The input of the adaptive filter is connected to the electric output of the hearing aid and the output signal of the adaptive filter may be subtracted from the input transducer signal to compensate for the acoustic feedback.
The hearing aid comprises an input transducer for transforming an acoustic input signal into a first electrical signal, a first filter bank with bandpass filters for dividing the first electrical signal into a set of bandpass filtered first electrical signals, a first set of combining nodes for receiving said set of bandpass filtered first electrical signals and combining them with a set of third electrical signals in order to output a first set of combining node output signals, a processor adapted for individual processing of each signal among the set of combining node output signals and adding together the processed electrical signals in order to generate a second electrical signal, an output transducer for transforming said second electrical signal into an acoustic output signal, a second filter bank with bandpass filters for dividing said second electric signal into a set of bandpass filtered second electrical signals, the bandpass filters of the second filter bank being substantially identical to respective bandpass filters of the filter bank, a first set of adaptive filters for estimating acoustic feedback by filtering of the bandpass filtered second electrical signals according to a set of the first filter coefficients and generating the set of third electrical signals, a second set of adaptive filters with second filter coefficients for filtering the bandpass filtered second electrical signals into respective fourth electrical signals, a second set of combining nodes for generation of fifth electrical signals by combining the fourth electrical signals with the respective signals of said first set of combining node output signals, and for inputting said fifth electrical signals to said second set of adaptive filters for adjustment of the second filter coefficients, and a controller adapted to determine a first parameter of an acoustic feedback loop of the hearing aid and to adjust a first adaptation rate of said set of first filter coefficients, wherein said controller is adapted to determine a second parameter of an acoustic feedback loop of the hearing aid and to adjust the second filter coefficients with a second adaptation rate that is higher than the first adaptation rate, and wherein said controller is adapted to estimate the amount of acoustic feedback based on information from said second set of adaptive filters.
It is an important advantage of the present invention that the requirement of a gain safety margin is significantly reduced since the controller automatically adjusts a parameter of the electronic feedback loop whenever the hearing aid operates with a high risk of generating undesired sounds so that such generation is substantially avoided.
In the following, the frequency ranges of the bandpass filters are also denoted channels.
The invention, in a second aspect, provides a hearing aid comprising an input transducer for transforming an acoustic input signal into a first electrical signal, a processor for generation of a second electrical signal by processing of said first electrical signals into a second electrical signal, an output transducer for transforming the second electrical signal into an acoustic output signal, a first adaptive filter with first filter coefficients for estimation of acoustic feedback by generation of a third electrical signal by filtering of said second electrical signal and adapting said third signal to said first first electrical signal, which first adaptive filter adaptive filter is a warped adaptive filter, wherein the first filter coefficients are updated with a first convergence rate, a set of second adaptive filters with second filter coefficients for filtering said second electrical signals into respective fourth electrical signals, and a combining node for generation of fifth electrical signals by combining the fourth electrical signals with the respective first electrical signals and for inputting the fifth electrical signals to said set of second adaptive filters, and wherein the second filter coefficients are updated with a second convergence rate that is higher than the first convergence rate.
The invention, in a third aspect, provides a hearing aid comprising an input transducer to transforming an acoustic input signal into a first electrical signal, a first filter bank with bandpass filters for dividing the first electrical signal into a set of bandpass filtered first electrical signals, a first set of combining nodes for receiving said set of bandpass filtered first electrical signals and combining them with a set of third electrical signals in order to output a first set of combining node output signals, a processor adapted for individual processing of each signal among the set of combining node output signals and adding together the processed electrical signals in order to generate a second electrical signal, an output transducer for transforming said second electrical signal into an acoustic output signal, a second filter bank with bandpass filters for dividing said second electrical signal into a set of bandpass filtered second electrical signals, the bandpass filters of the second filter bank being substantially identical to respective bandpass filters of the first filter bank, a first set of adaptive filters for estimating acoustic feedback by filtering of the bandpass filtered second electrical signals according to a set of first filter coefficients and generating the set of third electrical signals, and a controller adapted to determine an operating gain of the processor and to adjust a first adaptation rate of said set of first filter coefficients according to the operation gain.
In a simple embodiment of the invention, the hearing aid is a single channel hearing aid, i.e. the hearing aid processes incoming signals in one frequency band only. Thus, the first filter bank consists of a single bandpass filter, and the single bandpass filter may be constituted by the bandpass filter that is inherent in the electronic circuit, i.e. no special circuitry provides the bandpass filter. Correspondingly, the adding in the processor of processed electrical signals is reduced to the task of providing the single processed electrical signal at the output of the processor. Further, the second filter bank consists of a single bandpass filter, and the first set of adaptive filters consists of a single adaptive filter.
Typically, hearing defects vary as a function of frequency in a way that is different for each individual user. Thus, the processor is preferably divided into a plurality of channels so that individual frequency bands may be processed differently, e.g. amplified with different gains. Correspondingly, the hearing aid may comprise a first set of adaptive filters with a plurality of adaptive filters for individual filtering of signals in respective frequency bands whereby a capability of individually controlling acoustic feedback in each channel of the hearing aid is provided. Preferably, the frequency bands of the first set of adaptive filters are substantially identical to the frequency bands of the first filter bank so that the bandpass filters do not deteriorate the operation of the adaptive filters.
In one embodiment of the invention, the first set of adaptive filters subtracts the electrical output of the hearing aid from the input to the processor and the difference signal is used for modification of the filter coefficients as explained below. The difference signal is not used for modification of the input signal to the processor whereby distortion of the signal is avoided. Thus, in this embodiment of the invention the first adaptive filter is used for estimation of the acoustic feedback signal without distortion of the processed signal. Further, in this embodiment, at least one of the adaptive filters of the first set of adaptive filters may operate on a respective decimated bandpass filtered second electrical signal whereby signal processing power requirement is minimized without requiring additional filters since the adaptive filter output signal does not affect the processed signal directly.
In another embodiment of the invention, the first set of adaptive filters subtracts the electrical output of the hearing aid from the electrical signal from the input transducer and the difference signal is used for modification of the filter coefficients and is fed to the input of the processor whereby the acoustic feedback signal is substantially removed from the signal before processing by the processor. In this embodiment, decimation of signals may be employed in the processor and in the first set of adaptive filters if a third filter bank that is substantially identical to the first filter bank is added in the processor before summation of the individual processed signals from each processor channel to the output signal from the processor.
Generation of undesired sounds may be avoided by monitoring of the loop gain of the acoustic feedback loop, i.e. the gain of the acoustic feedback path from the output transducer to the input transducer including the transfer functions of the transducers plus the gain of the electronic circuitry included in the signal path from input to output of the hearing aid. When the loop gain approaches one, certain actions may be taken to prevent generation of unwanted sounds. Since the first set of adaptive filters generates a signal that corresponds to the signal generated by acoustic feedback, monitoring of attenuation in the first set of adaptive filters and of gains in corresponding channels of the processor provides an indication of the loop gain of the acoustic feedback loop. Thus, the controller may be adapted to monitor attenuation in the first set of adaptive filters, e.g. by determination of the individual ratios between the magnitude of the signal at the inputs of the individual filters and the signals at the corresponding outputs of the individual filters. Further, the controller may be adapted to monitor the gains of the individual channels of the processor, e.g. by a similar determination of input and output signal levels of individual processor channels, or by reading values from registers in the processor containing current gain values of individual processor channels. Typically, the processor channel gains are different for different channels and they are input level dependent.
Based on the monitoring of a first parameter of the acoustic feedback loop, such as the loop gain, the gain of a processor channel, the attenuation of an adaptive filter of the first set of adaptive filters, etc, a second parameter of the hearing aid may be adjusted to prevent generation of undesired sounds. For example, the gain of at least one processor channel may be modified, e.g. lowered, to keep the acoustic feedback loop gain below one.
The second parameter may be a maximum gain limit Gmax that the gain of the processor is not allowed to exceed within a specific channel. The adaptation rate of the first set of adaptive filters may be kept constant while the maximum gain limit Gmax of a specific channel of the processor is lowered whenever the hearing aid approaches a state in that channel with a high risk of generating undesired sounds, e.g. caused by a sudden change in the acoustic environment. For example, the maximum gain limit Gmax of a specific channel is lowered while the first adaptive filter adapts to a changed acoustic environment, and is restored to the original value when the adaptive filter has adapted to the new situation. Hereby, no distortion of the desired signal is generated.
It is an important advantage of this embodiment of the invention that the operating gain of the hearing aid may be very high without a risk of generating undesired sounds since the gain is automatically lowered if the feedback loop approaches resonance. Thus, a gain safety margin is substantially not required.
In embodiments wherein the bandpass filters of the second filter bank are substantially identical to respective bandpass filters of the first filter bank, each channel may be individually controlled based on a determination in that channel whereby reduction of gain by influence from frequencies outside the channel in question may be avoided.
Further, in an embodiment of the invention wherein the difference signal from the first adaptive filter is fed to the input of the processor, the second parameter may be a first convergence or adaptation rate of the first set of adaptive filters. For example, the adaptation rate of the filter may be made dependent on the operating processor gain in such a way that whenever the hearing aid approaches a state with a high risk of generating undesired sounds, e.g. caused by a sudden change in the acoustic environment, the adaptation rate of the first adaptive filter is increased to rapidly compensate for the change.
The convergence rate of the first set of adaptive filters may be adjusted by modifying the algorithm for updating the filter coefficients of the adaptive filter. As further described below, the algorithm may comprise one or more scaling factors that may be adjusted in response to the determination of the first parameter. For example, the one or more scaling factors may be adjusted as a predetermined function of the operating gains of the processor.
It is an important advantage of this embodiment that the operating gain of the hearing aid may be very high without a risk of generating undesired sounds since the closer the acoustic feedback loop gain approaches resonance the faster the adaptive filter will adapt to the situation. The fast adaptation of the adaptive filter may cause the desired signal to be distorted as previously described. However, as soon as the adaptive filter has adapted, the convergence rate is lowered and the desired signal is no longer distorted. Further, the distortion may take place in a frequency band that does not affect the intelligibility of the received sound signal.
A gain interval from G0 to Ga may be provided in the hearing aid. G0 is a predetermined lower gain limit below which feedback resonance and generation of undesired sounds can not occur. G0 may be determined during the fitting procedure. Ga is an adjustable upper gain limit that is adjusted according to desired sound quality. Preferably, Ga is adjusted during the fitting procedure.
The convergence rate may vary as a predetermined function, such as a linear or a non-linear function, of the gain of the processor, e.g. in the range from G0 to Ga. For example, one or more scaling factors of the updating algorithm of the adaptive filter may vary as a predetermined function, such as a linear or a non-linear function, of the gain of the processor, e.g. in the range from G0 to Ga.
During fitting of the hearing aid to the individual user, the transmission characteristics of the feedback path is measured. Based on these characteristics, the values of G0 and Ga with appropriate safety margins are determined and stored in the hearing aid. For determination of G0 there are several factors to take into consideration. The feedback path characteristics are, as already mentioned, not constant. Thus, sudden changes may lead to feedback resonance if the feedback compensation is too slow. Further, prediction of the magnitude and duration of changes of the attenuation of the feedback path may be difficult. On the other hand, fast adaptation may lead to unacceptable distortion of the desired signal, the level of unacceptable distortion again being a subjective quantity.
However, in situations where the characteristics of the acoustic feedback path have been stable for a certain period it is possible to estimate the characteristics of the feedback path accurately since in such a situation the relation between the signals at the inputs of the first set of adaptive filters and the signals at the outputs of the first set of adaptive filters is a precise measure for such characteristics, e.g. the attenuation, of the acoustic feedback path. Knowing the gain characteristics of the digital processor and of the acoustic feedback signal, an estimate for the acoustic feedback loop may be provided. From this knowledge, a dynamically changing value of G0 may be incorporated in the hearing aid. In one embodiment the interval from G0 to Ga may have a fixed size, independent of the changes in G0, i.e. the entire interval is shifted in accordance with changes of G0.
According to a preferred embodiment of the invention, the hearing aid further comprises a second set of adaptive filters operating in parallel with, i.e. on the same signals as, the first set of adaptive filters but with second convergence rates that are higher than the first convergence rates of the first set of adaptive filters. The outputs of the first set of adaptive filters are fed to the corresponding inputs of the processor whereby the acoustic feedback signal is substantially removed from the signal before processing by the processor. The outputs of the second set of adaptive filters are not used for modification of the processor input signals.
In this embodiment, the controller is adapted to estimate the amount of acoustic feedback by determination of a parameter of the second set of adaptive filters. The high second convergence rate allows the second adaptive filter to track the acoustic feedback more closely over time than the first adaptive filter. Further, since the output signal of the second adaptive filter is not subtracted from the input transducer signal, the desired signal is not distorted by the second adaptive filter.
Thus, according to a preferred embodiment of the invention, a hearing aid is provided further comprising a set of second adaptive filters with second filter coefficients for suppression of feedback in the hearing aid by filtering the bandpass filtered second electrical signals into respective fourth electrical signals, a combining node for generation of fifth electrical signals by subtraction of the fourth electrical signals from the respective bandpass filtered first electrical signals and for feeding the fifth electrical signals to the processor, and wherein the second filter coefficients are updated with a second convergence rate that is higher than the first convergence rate.
The amount of acoustic feedback may be estimated by determination of the ratio between the magnitude of the signals at the inputs of the second set of adaptive filters and the signals at the respective outputs of the second set of adaptive filters. This approach provides a quick response to changes in the acoustic feedback path and requires very little processor power.
The second parameter may be a second convergence or adaptation rate of the first set of adaptive filters. For example, the adaptation rate of the filtering may be made dependent on the operating gain of the processor or, the attenuation of the second set of adaptive filters or, a combination of the two, in such a way that whenever the hearing aid approaches a state with a high risk of generating undesired sounds, e.g. caused by a sudden change in the acoustic environment, the adaptation rate of the first adaptive filter is increased to rapidly compensate for the change.
As previously described for the second set of adaptive filters, the convergence rate of the first set of adaptive filters may be adjusted by modifying the algorithm for updating the filter coefficients of the adaptive filters. As further described below, the algorithm may comprise one or more scaling factors that may be adjusted in response to the determination of the second parameter. For example, the one or more scaling factors may be set as a predetermined function of the operating gains of the processor.
The first set of adaptive filters provides individual filtering of signals in respective frequency bands. Preferably, the frequency bands of the first set of adaptive filters are substantially identical to the frequency bands of the second filter bank.
The frequency bands of the first set of adaptive filters may differ in number and range from the frequency bands of the first filter bank and the second set of adaptive filters. However, in a preferred embodiment of the present invention, the second filter bank comprises a plurality of bandpass filters while the first set of adaptive filters consists of a single adaptive filter providing modification of the processor input signal in a single frequency band whereby a hearing aid with a frequency dependent hearing aid compensation capability is provided with a simple single band acoustic feedback compensation loop.
Thus, according to a preferred embodiment of the present invention, a hearing aid is provided further comprising a second adaptive filter with second filter coefficients for suppression of feedback in the hearing aid by filtering the second electrical signal into a fourth electrical signal, a combining node for generation of a fifth electrical signal by subtraction of the fourth electrical signal from the first electrical signal and for feeding the fifth electrical signal to the respective bandpass filters of the first filter bank, and wherein the second filter coefficients are updated with a second convergence rate that is higher than the first convergence rate.
Thus, in a preferred embodiment of the invention, the processor and the second adaptive filter are divided into channels covering the same frequency bands while the first adaptive filter is not divided into a plurality of channels. Further, the controller may be adapted to control the individual maximum gain limits Gmax of each processor channel in response to determination of the attenuation of the corresponding first adaptive filter channel. The controller may further be adapted to increase a second convergence rate of a filter of the second set of adaptive filters when the corresponding processor channel gain is limited by a Gmax limit so that the duration of the gain limitation may be decreased. Still further, the controller may be adapted to adjust the gain limit and/or the convergence rate in accordance with the current mode of operation of the hearing aid. The term mode of operation will be explained below.
Preferably, at least one adaptive filter is a finite impulse response (FIR) filter, and even more preferred at least one adaptive filter is a warped filter, such as a warped FIR filter, a warped infinite impulse response (IIR) filter, etc.
In the present example of a warped FIR filter, the unit delays are substituted by first order allpass sections. However, the warping may as well be realized with second order and even higher order allpass sections. A first order allpass section has the z-transform:             z              -        1              -    γ        1    -                  z                  -          1                    ⁢      γ      
where xcex3 is a warping parameter. Thus, the fixed delays in a FIR filter are substituted by frequency dependent delays leading to large delays at low frequencies and smaller delays at high frequencies. It should also be noted that the allpass elements are internally recursive and therefore warped FIR filters have infinite impulse responses. Thus, the term warped FIR is somewhat contradictory but describes well the structural analogy to transversal FIR filters.
In embodiments of the present invention, the order of a warped FIR filter may be considerably lower than the order of a FIR filter with comparable specifications. Thus, for a given circuit complexity, a warped FIR filter is capable of providing better filter characteristics than a FIR filter. Further, the warping parameter xcex3 may be used as a control parameter for controlling the transfer function, i.e. the positioning of resonances and cut-off frequencies in the frequency spectrum, whereby the spectrum of the error signal e(n), i.e. the difference between the filter output signal and the desired signal, may be minimized within a desired frequency range.
In the FIR or warped FIR filter, the next sample Y(t+T) is calculated according to the following equation:
Y(t+T)=c(t)u(t) 
wherein             c      _        ⁡          (      t      )        =                    (                                                            c                0                                                                                        c                1                                                                        ⋮                                                                          c                i                                                                        ⋮                                                                          c                                  N                  -                  1                                                                    )            ⁢              xe2x80x83            ⁢      and      ⁢              xe2x80x83            ⁢                        c          _                ⁡                  (          t          )                      =                  (                                                            u                0                                                                                        u                1                                                                        ⋮                                                                          u                i                                                                        ⋮                                                                          u                                  N                  -                  1                                                                    )            =              (                                                            u                ⁡                                  (                  t                  )                                                                                                        u                ⁡                                  (                                      t                    -                    T                                    )                                                                                        ⋮                                                                          u                ⁡                                  (                                      t                    -                    iT                                    )                                                                                        ⋮                                                                          u                ⁡                                  (                                      t                    +                    T                    -                    NT                                    )                                                                    )            
It is noted that u is an N dimensional vector containing the latest N samples of the signal u and c is a vector containing the N coefficients of the N""th order filter. T is the sampling period.
In the equation, u(t) is the actual value at the actual time t, and u(t-iT) is the signal value at i sampling periods prior to the actual time t. In discrete time systems, a shorthand notation is often used where the symbol u(i) indicates the signal value at the time t-iT, i.e. u(t-iT) in the equation above.
It is well known, e.g. cf. Adaptive Filtering by Paulo S. R. Diniz, Kluwer Academic Publishers, 1997, to use a least mean square algorithm for updating of the filter coefficients in an adaptive filter:
c(t+T)=c(t)+xcexcu(t)e(t) 
Using the above-mentioned shorthand notation (n is the reference number of the actual sample), the equation is rewritten:       (                                                      c              0                        ⁡                          (                              n                +                1                            )                                                                                      c              1                        ⁡                          (                              n                +                1                            )                                                            ⋮                                                                c              i                        ⁡                          (                              n                +                1                            )                                                            ⋮                                                                c                              N                -                1                                      ⁡                          (                              n                +                1                            )                                            )    ⁢      xe2x80x83    =            (                                                                  c                0                            ⁡                              (                n                )                                                                                                        c                1                            ⁡                              (                n                )                                                                          ⋮                                                                              c                i                            ⁡                              (                n                )                                                                          ⋮                                                                              c                                  N                  -                  1                                            ⁡                              (                n                )                                                        )        +          μ      ·              e        ⁡                  (          n          )                    ·              (                                                                              u                  0                                ⁡                                  (                  n                  )                                                                                                                          u                  1                                ⁡                                  (                  n                  )                                                                                        ⋮                                                                                            u                  i                                ⁡                                  (                  n                  )                                                                                        ⋮                                                                                            u                                      N                    -                    1                                                  ⁡                                  (                  n                  )                                                                    )            
Or in an even shorter form:
ci(n+1)=ci(n)+xcexcui(n)e(n) 
wherein i references the individual vector elements.
It is preferred to use a leaky least mean square algorithm for updating the filter coefficients:
ci(n+1)=xcex(ci(n)xe2x88x92ci(0))+ci(0)+xcexcui(n)e(n), 
where ui is a set of signal values derived from the output signal of digital processor in the n""th sampling period and the i-I preceding sampling periods, ci is a set of filter coefficients, e is the current value of the error signal and xcex and xcexc are scaling factors. The value of xcexc is typically in the magnitude of 10xe2x88x926 and the value of xcex is typically approximately 0.99. xcex is denoted leakage and when xcex less than 1, the filter coefficients will drift towards their respective initial values ci(0). xcexc is the convergence rate and determines the rate with which the adaptive filter adapts to a change. The adaptation rate increases with increasing values of xcexc.
It may further be advantageous to normalize the algorithm so that the adaptive filter, substantially, does not respond to momentary dynamic changes in the input signal. It should be noted that for the purpose of estimating the acoustic feedback signal, the desired input signal is irrelevant and constitutes noise deteriorating the convergence performance of the adaptive filter. The normalized algorithm is referred to as a normalized Least Mean Square (nLMS) algorithm:             c      _        ⁡          (              n        +        1            )        =            λ      ⁡              (                                            c              _                        ⁡                          (              n              )                                -                                    c              _                        ⁡                          (              0              )                                      )              +                  c        _            ⁡              (        0        )              +          μ      ⁢                                    u            _                    ⁡                      (            n            )                                                              u              _                        ⁡                          (              n              )                                ·                                    u              _                        ⁡                          (              n              )                                          ⁢                        e          ⁡                      (            n            )                          .            
However in the above equation the calculation of the power requires significant processing power and consequently, it is preferred to use a power estimate according to the equation:
Pu(t+T)=xcex1Pu(t)+(1xe2x88x92xcex1)u2(t) 
where xcex1 is a predetermined constant that determines the rate with which the Pu estimate changes. The algorithm is referred to as a power normalized Least Mean Square algorithm. The power estimate may also be based on the output signal from the input transducer so that the influence from sudden changes in the power of the input signal on the adaptation algorithm is minimized.
Further, a third update algorithm may be used for updating the adaptive filter coefficients denoted a leaky sign least mean square algorithm:
ci(n+1)=xcex(ci(n)xe2x88x92ci(0))+ci(0)+xcexcsui(n) 
where xcexcs is the sign of the e(n) signal multiplied by xcexc.
Still further, a fourth update algorithm that may be used for the adaptive filter coefficients denoted a leaky signxe2x80x94sign least mean square algorithm:
ci(n+1)=xcex(ci(n)xe2x88x92cl(0))+ci(0)+xcexcssgn(ui(n)) 
where sgn(ui(n)) is the sign of ui(n).
The filter coefficients may be updated based on a difference signal that is processed, e.g. combined with another signal, averaged or otherwise filtered, etc. Filtering may be performed in a focussed manner as known in the art.
Further, it should be noted that in a multichannel hearing aid according to the invention, the adaptive filters of the channels need not have identical number of taps. For example, it may be desirable to include more taps in adaptive filters operating in low-frequency channels.
As already mentioned, the controller may adjust xcex and xcexc in response to the determination of a first parameter of the acoustic feedback loop of the hearing aid.
Various sets of parameters of the hearing aid may be provided for various respective types of sound, e.g. speech, music, etc, that the user desires to hear and various respective types of acoustic environment, e.g. silence, noise, echo, crowd, open air, room, head set, etc, in which the user is situated. For example, various gain settings as a function of frequency may be provided, various gain settings as a function of input signal level may be provided, and various convergence rates as a function of operating processor gain may be provided, etc. Each set of parameters defines a specific mode of operation of the hearing aid and when the hearing aid operates with a specific set of parameters it is said to operate in the corresponding mode. Thus, in a specific mode of operation, specific parameter values of the hearing aid are set for appropriately processing of corresponding specific sounds in a specific acoustic environment. Likewise automatic adjustment of the parameters may be performed in accordance with the current mode of operation.
The type of sound may be selected by the user or, it may be automatically detected by the hearing aid, e.g. by a frequency analysis, analysis of signal to noise ratio at various frequencies, analysis of sound dynamics, speech recognition, recognition by neural networks, etc.
Likewise, the type of acoustic environment may be selected by the user or, it may be automatically detected by the hearing aid, e.g. by a frequency analysis, analysis of signal to noise ratio at various frequencies, analysis of sound dynamics, recognition by neural networks, etc.
For example, the user may desire to listen to music. The first convergence rate of the first adaptive filter may then be set to a value that is in conformance with the auto-correlation of music. Further, gain adjustments or adjustments of the first convergence rate may also be performed in conformance with the auto-correlation of music. For example, when the first convergence rate, e.g. one or more scaling factors, is controlled as a function of processor gain, the function may be selected from a set of functions, each of which is adapted for use in a specific acoustic environment with certain sounds, such as music, speech, etc, that the user has decided to listen to.
Furthermore, adjustments may also be performed in accordance with the rate of change of measured parameters, e.g. of the acoustic feedback path, e.g. the feedback gain, etc, etc.